The subject matter disclosed herein relates to computed tomography imaging systems and, more particularly, to an apparatus and method of x-ray voltage modulation for controlling patient exposure to radiation in such systems.
The imaging process in a computed tomography system utilizes an x-ray source and x-ray detectors to scan a patient, and to construct cross-sectional images of the patient from these scanned data. It is often desirable to limit the amount of patient exposure to x-ray radiation incurred during such scans. Referring to FIG. 1, a “third generation” computed tomography (CT) imaging system 10, for example, is configured to perform computed tomography imaging by means of photon counting and energy discrimination of x-rays at high flux rates. The CT imaging system 10 comprises a gantry 12, with a collimator assembly 18, a data acquisition system 32, and an x-ray source 14 disposed on the gantry 12 as shown. A support surface, such as a table 46, serves to move all or part of a target, such as a package (not shown) or a patient 22, through a gantry opening 48 in the gantry 12. During a scan to acquire x-ray projection data, the gantry 12 rotates about a center of rotation 24 along with the x-ray source 14 and the detector assembly 15.
Referring now also to FIG. 2, the x-ray source 14 projects a beam of x-rays 16 through the patient 22 onto a plurality of detector modules 20 in a detector assembly 12. The detector assembly 12 includes the collimator assembly 18, the detector modules 20, and the data acquisition system 32. Each detector module 20 in a conventional system produces an analog electrical signal that represents the intensity of an attenuated x-ray beam after it has passed through the patient 22. The data acquisition system 32 converts the sensed data to digital signals for subsequent processing. It can be appreciated that mage reconstruction for 3-D imaging requires a linear attenuation output. This may be achieved by initially performing a calibration at each of the kVp points of interest, typically using a water target, and can be a very time consuming operation.
Operation of the gantry 12 and the x-ray source 14 are controlled by a control mechanism 26. The control mechanism 26 includes an x-ray generator 28 that provides power and timing signals to the x-ray source 14, and a gantry motor controller 30 that controls the rotational speed and position (i.e., the gantry angle) of the moving components of the gantry 12. An image reconstruction processor 34 receives sampled and digitized x-ray data from the data acquisition system 32 and performs high speed reconstruction. Reconstructed images are applied as inputs to a computer 36 which can also store the images in a mass storage device 38.
The computer 36 also receives commands and scanning parameters input from an operator console 40. An associated display, such as a cathode ray tube display 42, allows an operator to observe the reconstructed image and other data from the computer 36. The commands and scanning parameters are used by the computer 36 to provide control signals and information to the data acquisition system 32, the x-ray generator 28, and the gantry motor controller 30. In addition, the computer 36 operates a table motor controller 44 which controls the motorized table 46.
Operating parameters for the x-ray source 14, such as peak x-ray tube kilovoltage (kVp) and x-ray tube current (mA), may be set from the operator console 40. To reduce patient dose, and to achieve improved image quality, the CT imaging system 10 may also have the capability of modulating the x-ray tube current during a scan. This modulation includes change to the x-ray tube current as the CT imaging system 10 scans at different height of the body, as well as x-ray tube current variation at the same patient height, but at different angular positions. The objective of x-ray tube current modulation is to deliver the optimal signal-to-noise ratio for each projection with improved patient dose management, such as by increasing x-ray tube current at longer mass penetration length. However, for a relatively fast scan speed, the operator may not be able to modulate the kVp to achieve the desired results, and the kVp may be left at a constant value. The prior art does disclose methods of modulating patient dose rate, such as exemplified in U.S. Pat. No. 6,233,310 “Exposure management and control system and method” and U.S. Pat. No. 7,545,915 “Dose rate control in an X-ray system,” but such procedures do not provide the accuracy and the output linearity required for 3-D imaging.
While the use of x-ray tube current modulation can provide for patient x-ray radiation dose reduction and may yield better image quality (IQ), certain shortcomings have been identified with this method. The tube current modulation speed in conventional x-ray tubes, for example, is limited by the thermal constant of the tube cathode and cannot be varied as rapidly as the tube current modulation method may require. Also, because of the speed with which the x-ray source 14 rotates around the patient 22, the system operator can not manually modulate the x-ray tube output within an individual rotation but can provide some voltage modulation for limiting patient exposure only as successive rotations of the x-ray source 14 occur. This shortcoming may become a more severe problem with newer, even fast-rotating x-ray scanners. Moreover, increasing the x-ray tube current, such as may be required for relatively long mass penetration lengths, may result in a linear signal increase at the detector, indicating linear dose contribution. In addition, a “photon starvation” problem may result at the highest x-ray tube current capability of an imaging system, for a given x-ray tube voltage.
What is needed is a method of controlling x-ray radiation exposure in imaging systems which addresses the shortcomings of the prior art.